The present invention relates generally to a scintillation material for making scintillation detectors. More specifically, the invention provides a scintillation material comprising nano-scale particles of a metal oxide, a metal oxyhalide, a metal oxysulfide, or a metal halide and methods for preparing the same.
Scintillators are materials that convert high-energy radiation, such as X-rays and gamma rays, into visible light. Scintillators are widely used in detection and non-invasive imaging technologies, such as imaging systems for medical and screening applications. In such systems, high-energy photons typically pass through the person or object undergoing imaging and, on the other side of the imaging volume, impact a scintillator associated with a light detection apparatus. The scintillator typically generates optical photons in response to the high-energy photon impacts. The optical photons may then be measured and quantified by the light detection apparatus, thereby providing a surrogate measure of the amount and location of high-energy radiation incident on the detector. Additionally, scintillators may be useful in systems used to detect radioactive objects, such as contraband or contaminants, which might otherwise be difficult to detect.
With regard to non-invasive imaging techniques, one of the most important applications for scintillators is in medical equipment for the production of radiographic images using digital detection and storage systems. For example, in current digital X-ray imaging systems, such as CT scanners, radiation from a source is directed toward a subject, typically a patient in a medical diagnostic application. A portion of the radiation passes through the patient and impacts a detector. The surface of the detector converts the radiation to light photons which are sensed. The detector is divided into a matrix of discrete picture elements, or pixels, and encodes output signals based upon the quantity or intensity of the radiation impacting each pixel. Because the radiation intensity is altered as the radiation passes through the patient, the images reconstructed based upon the output signals provide a projection of the patient's tissues similar to those available through conventional photographic film techniques.
Another high-energy radiation based imaging system is positron emission tomography (PET), which generally employs a scintillator-based detector having a plurality of pixels typically arranged in a circular array. Each such pixel comprises a scintillator cell coupled to a photomultiplier tube. In PET, a chemical tracer compound having a desired biological activity or affinity is labeled with a radioactive isotope that decays by emitting a positron. Subsequently, the emitted positron interacts with an electron giving out two 511 keV photons (gamma rays). The two gamma rays are emitted simultaneously and travel in opposite directions, penetrate the surrounding tissue, exit the patient's body, and become absorbed and recorded by the detector. By measuring the slight difference in arrival times of the two photons at the two points in the detector, the position of the positron inside the target can be calculated. The limitations of this time difference measurement are highly dependent on the stopping power, light output, and decay time of the scintillator material.
In both CT and PET, a small pixel size is required to generate an accurate image, i.e., for good spatial resolution. To avoid pixel to pixel contamination of the light produced in each luminescent module, the scintillators are made from single crystals or transparent ceramic imaging plates that are cut into small segments, or diced. The smaller segments are used with collimating reflectors between the individual elements to maintain as much of the light toward an individual detector as is physically possible. The dicing process limits the size of the individual pixel, as both production costs and process difficulties increase as the pixel size gets finer.
For systems where a still smaller pixel pitch is required, such as in digital radiographic systems, phosphors such as needles of CsI and fiber optic scintillator (FOS) face plates have been used. However these scintillators do not meet the more stringent luminescence requirements for CT systems. Scintillators based on CsI have a long decay time, leading to afterglow which tends to wash out images. Furthermore, detectors based on FOS plates do not have the high conversion efficiency needed for accurate imaging.
In contrast to the complex scintillators used for imaging applications, scintillators used in the detection of radioactive contraband or contamination are often simple plastic films, made from such materials as polythiophene or polyaniline. However, these systems are not very specific to the type of radiation involved, and often may give false alarms.
Accordingly, there is a need for new scintillators that can be easily formed into materials with the small pixel sizes needed for application in CT and PET, while affording transparency and tailored luminescence properties.